Over half of all patients in the US with cancer are treated with radiation therapy.
Radiation therapy is based on irradiating the patient, more particularly his or her tumor, with ionizing radiation. In the particular case of proton radiation therapy, the radiation is performed using a proton beam. It is the dose of radiation delivered to the tumor which is responsible for its destruction. Proton therapy is a desired form of radiation therapy because, in comparison to standard x-ray radiation therapy, proton therapy allows an increased dose of radiation to a tumor while reducing the amount of radiation to the healthy tissue surrounding it.
The central challenge to modern radiation therapy is to enhance local tumor control using dose escalation and to minimize the dose to normal tissues in order to improve survival and the quality of life of the patients. Radiation can damage normal tissue and thus causes both short term and later stage tumors in long-term cancer survivors.
The recent progress made by 3D conformal and intensity-modulated radiation therapy has reduced short term radiation-induced complications especially in dose-limiting organs like the brain, lung and intestine. Yet, acute short term complications do occur and are still the limiting factor for some treatments. More insidious for younger patients is the long term potential occurrence of secondary tumors for years and even decades after the treatment.
The replacement of x-ray therapy with protons will have substantial long term benefit to patients due to greatly reduced long and short term toxicity side effects. Such effects also have substantial costs associated with their treatment that may continue for many years after treatment. If the cost of proton therapy can be made about equivalent to that of x-rays for the primary treatment, they would result in substantial long term savings to health care providers.
One of the major roadblocks to the greater application of proton therapy is that of cost. The capital cost of a proton therapy center is due both to the cost of the very large equipment and of the heavily shielded vaults in which it is installed. However, the technology now included in FDA approved proton systems dates from some 20 years ago. It does not reflect advances made by researchers in the technology for use in research laboratories, in particular the use of high temperature superconductors.
A conventional proton therapy facility is typically composed of the particle accelerator, the proton beam guiding and controlling device, and the treatment rooms. The particle accelerator actually is not the dominating component in the layout of a conventional proton therapy facility where the accelerated beam is shared by several treatment rooms. When one takes a bird's eye view of a traditional four room facility, it is evident that the largest component of the overall proton beam delivery system is the proton beam guiding and controlling device, hereafter referred to as the gantry. The gantry transports and delivers the proton beam into a treatment room, bends the beam until it can be incident orthogonal to the patient and then rotates the beam around the patient. A typical gantry is comprised of large magnets, an evacuated pipe, a nozzle and a counterweight. All of the components are mounted on a large steel beam “squirrel cage” to enable the rotation of the proton beam around the patient.
FIG. 1 is one embodiment of a conventional treatment room floor layout in two dimensions. The gantry 10 is installed in the room shown at the top of the diagram. FIG. 1 shows the dominance of the gantry size on the overall layout of an extant proton therapy facility. The approximate weight of the convention proton gantry shown in FIG. 1 is approximately 120 tons. It is even more evident if one looks at all 3 dimensions and takes into account that the diameter of the gantries exceed 13 meters. Conventional isocentric gantries are generally greater than 13 meters in diameter and up to 15 meters long.
Another feature that drives the size of extant gantries is the need to have a very long nozzle, typically over 3 meters in length, that allows for an advanced scanning technique called spot beam scanning, also known as pencil beam scanning, with minimization of the entry dose at the skin. (A nozzle of at least 2.3 meters is required even if scanning is not performed.) The spot beam scanning technique spreads a small diameter incident proton beam over the target area at a certain depth inside the patient. Electromagnets mounted in a nozzle sweep the beam in two dimensions, X-Y, over the target area. In addition the beam intensity for each 3 dimensional spot (voxel) of the target area is varied to achieve a dose distribution that conforms exactly to the target area at a precise depth. Repeating this process for a range of decreasing energies (energy stacking) allows treatment at different depth and hence of the full tumor volume with any arbitrary shape. In one extant system beam scanning is performed in one direction upstream of the last 90 degree bending magnet of the gantry. While beam scanning upstream of the last gantry dipole did effectively reduce the treatment nozzle it also resulted in a very large, heavy gantry and included a costly 90 degree bend magnet weighing about 90 tons.
Conventional gantries now on the market capable of doing advanced beam scanning techniques such as spot scanning cost in the range of $10 million to $15 million installed and are mostly over 40 feet in diameter.
Superconducting technology has been proposed for ion therapy applications before, especially for carbon ion therapy. However, at that time it was concluded that complexity of the cryogenic system and difficulty of ramping the superconducting magnets would make the application of low temperature high field magnets technically overly challenging. Selecting conventional room temperature magnets for the carbon ion gantry design at Heidelberg facility resulted in a gigantic 630 tons structure that is 16 meters in height.
An unconventional design was developed at the Paul Scherrer Institute, where the X-Y beam scanning is performed in one direction upstream of the last 90 degree bending magnet of the gantry. However, the Paul Scherrer Institute design is not isocentric around the patient and consequently it also was necessary to rotate the patient in a concentric circle that at times left the patient over 2 meters above the ground. Nevertheless, beam scanning upstream of the last gantry dipole did effectively reduce the gantry diameter from 12 meters to about 5 meters for a gantry based on conventional room temperature dipoles. Unfortunately, this also resulted in a very large, heavy (about 90 tons), and costly 90 degree bend magnet.
A superconducting gantry design recently has been proposed for particle therapy facilities. The gantry design is based on Fixed Field Alternating Gradient (FFAG) magnets, which reduce the total weight of the gantry. However, such a gantry remains very large. The gantry measures about 20 meters in length, from the rotation point to the isocenter, and has a height of about 3.2 meters. Additionally, the gantry requires a significant increase in the number of magnetic elements in the gantry system.
Thus, there is a need in the art for development of new concepts that will greatly reduce the size and associated costs of the dose delivery gantry. It is an object of the present invention to reduce the installed cost of the gantry to about 33% of that of the current offerings, the size by about 50% and the weight by a factor of ten.